In X-ray-based imaging, X-ray emitters may be used as a particle source. Often, the X-ray emitters are pulsed in order, for example, to keep the patient dose low during medical imaging. The configuration of a known X-ray emitter or X-ray tube is depicted in FIG. 1.
FIG. 1 depicts a cross-section of an X-ray emitter. In a vacuum tube 1, electrons 3 are released into the vacuum by heating an emitter 2 (e.g. cathode). The electrons 3 are accelerated towards the anode 4 by a high voltage, that is applied between the emitter 2 and an anode 4. Upon striking the anode 4, approximately 1% of the energy of the electrons 3 is converted into X-ray radiation 5, the remaining energy transitioning into heat. The generated heat must be continuously extracted from the anode 4, otherwise there is a risk of the focal path melting on the anode 4. Therefore, in heavy-duty tubes, rotating anodes may be used in combination with directly cooled bearings. The X-ray radiation 5 exits the vacuum tube 1 through the outlet window 6. Using the motorized drive 7, the anode 4 is set in rotation.
For angiography, for example, a pulsed X-ray radiation is used to reduce the X-ray radiation exposure of the patient. The pulse generation for X-ray tubes may be realized by two methods.
A simple method is to switch the high voltage applied between the emitter 2 and the anode 4 on and off (primary pulsed X-ray radiation 5). Alternatively, a grating 8 may be arranged between the emitter 2 and the anode 4, or a metal element may be arranged around the emitter 2, that is exposed to a pulsed blocking voltage, and the electrons 3 are shielded toward the high voltage field. The grating 8 is switched on and off. The grating 8 alternately blocks and guides secondary pulsed X-ray radiation.
The first method is simple to realize. However, high capacities through cables etc. include disadvantage when actuating that the high capacities must be discharged via the X-ray tube causing electrons 3 to briefly strike the anode 4 with a lower energy, having experienced a lower acceleration voltage. The lower-energy electrons 3, however, generate a lower-energy X-ray radiation 5 that, in medical imaging for example, causes unnecessarily high patient doses without contributing to the imaging.
The second method creates clean X-ray pulses limited in terms of time, since the tube current are shielded in a very short switching time. A grating arrangement is disclosed by way of example in patent application DE 10 2009 004 186 A1. In heavy-duty X-ray tubes, however, the tube currents are very high. In order to generate such high tube currents, large emission surfaces of the emitter 2 are helpful. Large emission surfaces, however, are difficult to block. To this extent, a secondary pulsing is difficult in X-ray tubes with large currents.